REVIEW PAPER

Nanobiosensors: optofluidic, electrical and mechanical approachesto biomolecular detection at the nanoscale

David Erickson Æ Sudeep Mandal Æ Allen H. J. Yang ÆBernardo Cordovez

Received: 1 March 2007 / Accepted: 2 July 2007 / Published online: 8 August 2007

� Springer-Verlag 2007

Abstract Next generation biosensor platforms will

require significant improvements in sensitivity, specificity

and parallelity in order to meet the future needs of a variety

of fields ranging from in vitro medical diagnostics,

pharmaceutical discovery and pathogen detection. Nano-

biosensors, which exploit some fundamental nanoscopic

effect in order to detect a specific biomolecular interaction,

have now been developed to a point where it is possible to

determine in what cases their inherent advantages over

traditional techniques (such as nucleic acid microarrays)

more than offset the added complexity and cost involved

constructing and assembling the devices. In this paper we

will review the state of the art in nanoscale biosensor

technologies, focusing primarily on optofluidic type devi-

ces but also covering those which exploit fundamental

mechanical and electrical transduction mechanisms. A

detailed overview of next generation requirements is pre-

sented yielding a series of metrics (namely limit of

detection, multiplexibility, measurement limitations, and

ease of fabrication/assembly) against which the various

technologies are evaluated. Concluding remarks regarding

the likely technological impact of some of the promising

technologies are also provided.

Keywords Nanobiosensors � Biosensors � Optofluidics �Nanotechnology � Photonic crystal �Surface plasmon resonance � Nanowire � Cantilever

1 Introduction

Interest in the development of new biosensing and high-

throughput screening technologies has been largely driven

by recent developments in two different application areas.

The first of these areas grew out of a number of advance-

ments in proteomics and genomics research that led to an

explosion in the number of biomarkers associated with

specific disease states (Growdon 1999; Sander 2000;

Srinivas et al. 2001, 2002) and pharmacological responses

(Ross et al. 2004). While in general the number of useful,

unequivocal single biomarkers available for screening

healthy populations for complex diseases remains rela-

tively small (Ward et al. 2001; Hernandez and Thompson

2004) the potential for diagnosis based on panels of mul-

tiple biomarkers (Sidransky 2002; Wulfkuhle et al. 2003)

presents an exciting possibility. The second broad appli-

cation area relates to the need to rapidly detect and identify

emerging pathogenic threats. Viral detection [e.g. influenza

(Dawson et al. 2006; Liu et al. 2006; Lu 2006), adenovirus

(McCaman et al. 2001; Gu et al. 2003) and dengue hem-

orrhagic fever (Guzman and Kouri 2004; Zaytseva et al.

2005)] necessitates the development of sensor platforms

with very low limits of detection and false alarm rates that

can also track or at least account for the relatively high

rates of mutagenic drift. Other threats such as water-

(Straub and Chandler 2003), food- (Rasooly and Herold

2006) or air-borne pathogens (McBride et al. 2003; Stet-

zenbach et al. 2004) also require high sensitivity and

specificity but introduce other engineering challenges such

D. Erickson (&) � B. Cordovez

Sibley School of Mechanical and Aerospace Engineering,

Cornell University, 240 Upson Hall, Ithaca, NY 14853, USA

e-mail: de54@cornell.edu

S. Mandal

Applied and Engineering Physics, Cornell University,

Ithaca, NY, USA

A. H. J. Yang

Chemical and Biomolecular Engineering,

Cornell University, Ithaca, NY, USA

123

Microfluid Nanofluid (2008) 4:33–52

DOI 10.1007/s10404-007-0198-8

as larger sample volumes to be processed or more complex

target capture requirements. Emerging autonomous and

networked biosensor systems, which are designed to

operate and report findings with as little human interven-

tion as possible, add additional constrains such as power

usage and time multiplexing.

In this article we review the state of the art in nanoscale

biosensing techniques with a particular focus on optofluidic

approaches (i.e. those which exploit some combination of

optics and micro-/nanofluidics) but also inclusive of nano-

scopic electrical and mechanical devices. The goal will be to

introduce the various approaches in the context of how well

they will be able to address the challenges associated with the

two application areas described above. In the following

section a detailed overview of ‘‘next generation’’ biosensor

requirements is presented along with a final section which

describes where state of the art technologies tend to be

lacking and where these emerging technologies are likely to

have the biggest impact. Following this a detailed overview

of the various competing formats are presented. This review

is written assuming some knowledge of traditional biosen-

sing techniques. For more in-depth information on

biosensing, readers are referred to recent books by Kress-

Rogers (1997) or Eggins (1996).

2 Overview of next generation biosensor requirements

Though the application areas listed in the introduction are

reasonably disparate, in general the requirements for next

generation biosensor and biosensor platforms can broadly

be classified into three different categories. Here we pro-

vide a detailed overview of these requirements to provide

context for our later discussion.

2.1 Sensitivity and specificity

Biosensor development has always attempted to push the

limits of sensitivity and specificity with the goal of detecting

rarer targets with greater precision. The term sensitivity is

used in the literature in a number of different contexts

(D’Amico and Di Natale 2001) and thus it is useful to

delineate them now for clarity. Test sensitivity refers to the

number of true positives a given test yields (i.e. a given test

suggests the presence of a particular biomarker and its

presence can be confirmed by a ‘‘gold standard’’ test) divided

by the number of true positives and false negatives (i.e. the

test suggests the presence of a given biomarker and it is not

confirmed by the gold standard test). For any final technol-

ogy it is ultimately this measure that is of utmost importance.

As a metric which can lead to improved test sensitivity,

internal sensitivity is essentially the local slope of the sensor

response curve. In the case of the linear sensor response

curve this reduces to change in sensor output per unit change

in the quantity measured. As an example an appropriate

measure for wavelength sensitive optically resonant refrac-

tive index sensor would be Dk/RIU [i.e. the total expected

change in the resonant wavelength per unit change in

refractive index unit (RIU)]. The third measure of sensitivity

which is often encountered in the literature is the limit of

detection, LOD, or the smallest amount of a quantity of

interest which produces a measurable output signal. This

final measure is of particular interest since the lower the LOD

the earlier a disease state can be diagnosed or the presence of

an unexpected pathogen can be detected. Rather than simply

expressing the LOD as an absolute quantity (ng) it is com-

mon to divide this number by the volume of sample which is

processed (ng/mL) since this is also a the parameter of

interest. In general, however, the latter of these is dependent

on a much larger number of system parameters beyond just

that of the sensor performance (e.g. residence time, transport

speed, mixing efficiency, diffusivity of the target) thus

making it more difficult to provide a direct comparison. As

such in this review we will focus on the former.

2.2 Multiplexing

New sensor platforms must have inherently high degrees of

multiplexing capabilities. The necessity to maximize the

number of biomarkers a patient sample can be interrogated

against is perhaps obvious (i.e. the greater the number of

disease markers which can be probed, the more information

that can be obtained). In pathogen detection multiplexing

capability is required not only to screen against the large

number of different pathogens which may be present in a

given sample but also to provide specific subtype informa-

tion. In virology for example, subtype information is

particularly important for tracking emerging viruses (Daw-

son et al. 2006) and designing/delivering appropriate

vaccines. In cases where tracking a single biomarker is of

interest it may be also useful to provide time multiplexing

(i.e. samples taken at regular time intervals to be interrogated

against the same set of probes but at different sites so as to

avoid cross concentration). This allows for time tracking of,

for example viral load, against the individual base state,

enabling more accurate diagnosis. This time multiplexing

capability is more relevant to the development of autono-

mous sensor systems, where samples taken at regular

intervals.

2.3 Reduction in measurement complexity and cost

In general sensor systems which can reduce the number of

required sample processing steps as well as the amount of

34 Microfluid Nanofluid (2008) 4:33–52

123

on-chip or off chip infrastructure are likely to be more

successful than those which do not. Related to the former

of these is the goal of ‘‘label free’’ transduction methods

whereby the target itself does not require tagging with, for

example, a florescent label. Label free methods are typi-

cally based on a positive binding event inducing a change

in the local refractive index (which is most relevant for

optofluidic technologies), electrical, or mechanical condi-

tions and form the basis of most of the techniques

described herein. Techniques which successfully decrease

the LOD can also serve to eliminate some processing steps

such as target amplification. Nucleic acid amplification via

the polymerase chain, PCR, requires the sample to be sent

through a predefined thermal cycle each time roughly

doubling the amount of the target sequence. Although

numerous rapid chip based PCR devices have been dem-

onstrated (Kopp et al. 1998; Zhang et al. 2006a, b) reliance

on such processes limits the degree of multiplexing, which

can be achieved and requires more on-chip infrastructure.

In some portable cases the amount of power required to

perform the cycling is also detrimental. At present, most

technologies require relatively sophisticated forms of

infrastructure to perform the desired measurement. As such

the development of techniques that can couple high sen-

sitivity with a simple (e.g. colorimetric) feedback

technique would also be desirable. With regards to cost,

biosensor elements should minimize any cost prohibitive

fabrication or extensive assembly processes that would

make mass-producibility unfeasible. Although many

applications will be able to support higher device costs if

they can meet stringent sensitivity/specificity/autonomy

requirements, others, such as large scale screening efforts,

it is likely that cost will be an extremely important factor.

2.4 How well do existing technologies meet

these needs?

With respect to these requirements, the strength of con-

ventional array based technologies (e.g. nucleic acid and

protein microarrays or traditional ELISA approaches) has

always been in addressing the multiplexing requirements.

The incorporation of a suitable microfluidic element can

also serve to solve some of the requirements associated

with the third category above (Sect. 2.3). In general,

however, the relatively low sensitivity of such approaches

limits the types of targets which can be reasonably

expected to be interrogated and places stricter requirements

on the amount of sample processing and detection infra-

structure that is required. Emerging nanotechnologies such

as nanoparticles (Seydack 2005), nanowires and nanotubes

(Li et al. 2005; Zheng et al. 2005), nanomechanical reso-

nators (Majumdar 2002) and nanophotonics (Chow et al.

2004; Schmidt et al. 2004; Ouyang et al. 2006) are of

interest largely to address this failing. As will be described

in detail below, while many of these devices tend to not to

have much greater internal sensitivity (i.e. slope of the

sensor response curve in response to changes in bulk

properties such as conductivity or refractive index) than

traditional techniques, their inherent advantage is that the

total surface area or volume that is probed tends to be much

smaller. As a result the total mass required to impart a

measurable transduction signal is significantly lower and

the LOD is significantly improved. In general, however, the

extension of these technologies to the extreme parallelity of

the 2D microarray format is complicated by the challenges

involved in functionalization of individual sensor and 2D

optical or electrical addressing of reaction sites with sub-

micrometer spacing. The proceeding sections will focus

largely on how well these emerging nanotechnologies have

done in terms of pushing the limits of detection and what

approaches are being used to achieve the cost and paral-

lelity afforded by the conventional techniques.

3 Optofluidic biosensors

This first section reviews the state of the art in label free

optical biosensing techniques. This section is divided into

three areas. The first area focuses on techniques that exploit

localized changes in the refractive index, induced by bio-

molecular binding, in the evanescent field of a dielectric

structure. The second area is an extension of the first

looking more specifically at photonic crystal based devices.

The final section provides a detailed review of surface

plasmon resonance (SPR) based biomolecular detection.

3.1 Evanescent field based devices

While most of the optical energy is confined within the

structure itself, solid core dielectric waveguides have an

exponentially decaying tail of the guided optical mode,

referred to as the evanescent field, that impinges a small

distance (on the order of a hundred nanometers for most

systems of interest here) into the surrounding medium. The

majority of label-free optofluidic biosensors utilize this

evanescent field to probe the surface of a sensing site for

the presence of bound or absorbed analytes (Marazuela and

Moreno-Bondi 2002). Binding of the target at the sensing

site causes a change in the local refractive index in that

region imparting a slight phase shift to the propagating

optical mode. This simple phenomenon can be exploited

via a number of different techniques to perform label-free

optofluidic detection. In this first sub-section we present an

overview of these techniques while analyzing their

Microfluid Nanofluid (2008) 4:33–52 35

123

advantages and limitations with regards to the criteria

established above.

3.1.1 Interferometric based techniques

Interferometry can be used to detect the phase difference

between two collimated light beams of a coherent light

source. The simplest practical configuration for chip based

integrated optofluidic biosensing is the Mach-Zehnder

interferometer (MZI) such as that presented by Luff et al.

(1998). Their design consisted of an input optical wave-

guide, which splits into two arms of equal length and then

recombined to form the output optical waveguide. One arm

is referred to as the reference arm while a section of the

other, sensing arm, is functionalized with the desired bio-

recognition agent. In the absence of any surface

modifications to either of the arms, the light recombining at

the output port remains in phase giving rise to constructive

interference and maximal light intensity at the output port.

When binding occurs at the surface of the sensing arm it

changes the local refractive index and the resulting phase

shift causes the output power to drop due to destructive

interference effects. In the aforementioned work by Luff

et al., this phase shift was deduced by measuring the light

intensity at the output waveguide. More recently an MZI

biosensor using silicon nitride waveguides and standard

CMOS processes was demonstrated (Prieto et al. 2003). It

was used to detect the pesticide Carbaryl and was shown to

have a refractive index LOD of around 10�5. Others

(Heideman and Lambeck 1999) have managed to push this

limit to as low as 10�7 using the MZI configuration. While

this technique is very sensitive, it is difficult to multiplex

devices in a MZI configuration since each can consist of

only one reference and one sensing arm. Another drawback

is that the interaction length necessary for producing suf-

ficient phase shifts is often on the order of a centimeter,

which in comparison with some of the other techniques that

will be discussed below, is very large. In a variant on the

traditional approach, Lou et al. (2005) recently proposed

using silica nanowire waveguides in a MZI configuration.

Their theoretical estimates indicate that the sensing site

length can be decreased by an order of magnitude (to

around 1 mm) to achieve similar refractive index sensi-

tivity as the previously described works.

Brandenburg (1997) demonstrated the use a Young’s

interferometer based sensor for refractive index measure-

ments in liquids. Recently Ymeti et al. (2007) extended this

concept to perform multiplexed label-free biosensing using

the architecture shown in Fig. 1. These devices consist of a

waveguide which guides monochromatic light coupled in

from a laser source. This waveguide is split into two par-

allel waveguides by the means of a Y-junction. Similar to

the MZI sensor, one of the waveguides is functionalized

with a biorecognition agent while the other acts as a ref-

erence waveguide. The light emerging from the output ends

of the two waveguides is collected by a cylindrical lens and

made incident on a charge coupled device (CCD) screen.

By analyzing the interference pattern generated, the phase

shift caused by binding events occurring on the surface of

the sensing waveguide can be inferred. In their design they

were able to resolve the interference patterns between four

waveguides. Thus by using one waveguide as the reference,

three different targets could be detected simultaneously. In

that work detection of herpes simplex virus 1 (HSV-1) was

demonstrated and a device sensitivity of 10�7 RIUs, which

corresponded to a protein mass coverage of approximately

20 fg/mm2 was reported. Although a certain level of mul-

tiplexing has been demonstrated, it is very hard to push

beyond this limit with the Young interferometer design. As

in the case of MZI sensors, interaction lengths in these

sensors are on the order of a centimeter.

3.1.2 Resonant cavity based techniques

As stated above, one of the chief drawbacks of evanescent-

wave sensing by interferometric techniques is the long

interaction length of the sensing site, which requires a

relatively large amount of bound mass to make an appre-

ciable difference in the transduction signal. Resonant

cavity sensors provide a way to overcome this disadvan-

tage by shrinking the size of devices by orders of

magnitude while still retaining similar device sensitivity.

Microcavities sustaining whispering gallery modes

(WGMs) (Ilchenko and Matsko 2006; Matsko and Ilchenko

2006) have been a popular choice for a label-free biosen-

sing architecture. WGMs correspond to light being

confined along a circular orbit along the edge of a sphere,

disk or cylinder type structure. Those wavelengths of light,

which after completing one revolution return in phase, are

sustained in the resonator while the other wavelengths die

out due to interference effects. WGMs have been exten-

sively studied in liquid droplets and fused silica spheres

(Arnold et al. 2003) both of which can have nearly atomic

scale smoothness. In such microcavities optical losses are

significantly lower than in other optical resonators and the

Quality-factor (Q-factor) can exceed a hundred million

(Gorodetsky et al. 1996; Armani et al. 2003). Typically,

light is evanescently coupled into these resonators using

tapered fibers. The output spectrum observed at the end of

the coupling fiber consists of a series of sharply peaked

dips in transmission. Changes in the local refractive index

at the surface of the resonator cause a slight perturbation to

the resonance condition for the cavity, imparting a lateral

shift to the peaks in the output spectrum.

36 Microfluid Nanofluid (2008) 4:33–52

123

This concept was used to demonstrate a highly sensitive

refractometric sensor using fused silica microsphere reso-

nators (Hanumegowda et al. 2005). The microspheres had a

radius of 55 lm and the sensor had a LOD on the order of

10�7 RIU. For biosensing purposes these resonators can be

coated with a suitable biorecognition agent. Vollmer et al.

(2002) demonstrated detection of bovine serum albumin

(BSA) and streptavidin-biotin binding using such a design.

Although these systems possess the properties to make

them excellent candidates for ultra-sensitive biosensors, it

can be difficult to control the physical parameters of these

structures during fabrication as traditional lithographic

techniques typically cannot be used to create on-chip

cavities with such high Q-factors. This problem was solved

by Armani et al. (2003) who used a novel process to fab-

ricate toroidal silica microresonators with a diameter of

120 lm on a chip. These structures had a Q-factor in

excess of a hundred million. An extension of this work

(Armani and Vahala 2006) demonstrated the use of these

microresonators to detect heavy water concentration. Using

an entirely different approach, Zhu et al. (2007) have

proposed a technique for on-capillary refractive index

detections. The circular cross section of the capillary acts

as a ring resonator along which the input laser light remains

confined. The interaction of the evanescent field with the

contents of the liquid filled capillary allowed for non-

invasive, on-capillary analysis.

While the techniques described above are well suited to

perform highly sensitive detections, they lack robustness

and are difficult to integrate in planar systems compatible

with traditional microfluidics for performing multiplexed

detections. Another drawback is that the entire surface of

such devices has to be functionalized although the WGM

interacts with a small fraction of this surface area. Planar

microdisk (Boyd and Heebner 2001; Krioukov et al. 2002)

and microring resonator structures can overcome some of

these difficulties as they can easily be integrated on chip

using standard semiconductor fabrication techniques

(though the Q-factor is not nearly as high as those dis-

cussed above). Microring resonators consist of a ring

waveguide which is adjacent to a bus waveguide. Light

from a laser is coupled into the bus waveguide and this in

turn evanescently couples into the ring resonator. As in the

case of the silica microsphere resonators, resonance occurs

for those wavelengths which are in phase after performing

one round-trip around the ring and the spectrum at the

output end of the bus waveguide consists of sharply peaked

dips in transmission. Binding events along the surface of

the microring increase the refractive index in the evanes-

cent field, effectively lengthening the ring and causing the

resonant peaks to red shift. Chao et al. (2006) demonstrated

polymer microring resonators of 45 lm radius having a Q-

factor of 20,000. They were able to detect an effective

refractive index change of 10�7 RIU and had a detection

limit of approximately 250 pg/mm2 mass coverage on the

microring surface. While microring resonator sensors pro-

vide a robust architecture for potentially building highly

multiplexed biosensors, their binding surface area is still

reasonably large. For example in the Chao et al. device

mentioned above the total surface area or the ring resonator

is *6.5 · 10�4 mm2 thus the mass LOD is *160 fg.

3.2 Photonic crystal devices

Photonic crystals (Joannopoulos et al. 1995) are composed

of periodic dielectric structures. One of the features this

periodicity gives rise to is a range of wavelengths which

Fig. 1 Four-channel Young’s

interferometer based optical

biosensor from Ymeti et al.(2007). Channels 1, 2, and 3 are

the sensor channels which are

functionalized with antibody

capture probes and 4 is the

reference channel. Copyright

American Chemical Society.

Reproduced with permission

Microfluid Nanofluid (2008) 4:33–52 37

123

are not allowed to propagate within the structure, referred

to as the photonic bandgap. The size of the bandgap and its

position in the spectrum can be tuned by varying the

refractive index contrast of the dielectric materials and/or

the periodicity of the structure (Erickson et al. 2006). These

properties of photonic crystals make them extremely useful

in a number of applications, including biosensing. As an

example, Skivesen et al. (2007) demonstrated a photonic-

crystal waveguide biosensor. This consisted of a silicon

waveguide flanked on either side by a 2D photonic crystal

which caused light corresponding to the photonic bandgap

to remain guided in the waveguide. As with the other

optical devices, adsorption of proteins on the surface of the

photonic crystal increased the local refractive index and

shifted the bandgap. This was then detected by recording

the spectrum at the waveguide output. A total of 0.15 lM

concentration of BSA, corresponding to a surface coverage

of 6 ng/mm2, was easily detected. While being a novel

technique for performing label-free sensing, this device

design is hard to multiplex since the bandgap prohibits

transmission over a large range of wavelengths. In the

experiment mentioned above, the device exhibited a rela-

tively low solution phase LOD although the authors argue

that there is still room for optimization.

Photonic crystal resonator devices posses very high Q-

factors and are very sensitive to changes in the refractive

index of their structural elements. They consist of a 1D or

2D photonic crystal with a defect in the crystal structure

which acts as the resonant cavity. Chow et al. (2004)

demonstrated a 2D photonic crystal microcavity consisting

of a periodic lattice of holes in a silicon layer with a central

hole defect. Changes in the refractive index in these holes

shifted the resonant peak which allowed them to measure

the refractive index of the surrounding liquid medium.

More recently Lee and Fauchet (2007) demonstrated a

similar architecture with a Q-factor well over 5,000. The

device and experimental setup is shown in Fig. 2. They

were able to detect BSA with a lower detection limit of a

molecule monolayer with a total mass as small as 2.5 fg.

Although it is difficult to perform multiplexed detections

on a single waveguide using this design, it offers the unique

advantage of performing highly sensitive detections in

ultra-small volumes.

3.3 Surface plasmon resonance biosensing

Surface plasmons are electromagnetic waves which prop-

agate along metal/dielectric interfaces. As will be

described below the conditions for exciting these optical

modes are extremely sensitive to the dielectric environment

very near this interface. As a result SPR is one of the most

commonly exploited label-free optical biosensing tech-

nique in use today and forms the third optofluidic

architecture we will review here. Detailed reviews of the

topic are widely available in the literature (Homola et al.

1999a, b; Haes and Van Duyne 2004; Karlsson 2004) and

thus here we provide just a brief introduction to the various

implementations and focus on the limitations of the current

state of the art in SPR biosensing in the context of multi-

plexing and sensitivity.

3.3.1 Angular SPR biosensing

The most common implementation of SPR biosensing is

the attenuated total reflection (ATR) approach using the

Kretschmann geometry shown in Fig. 3a. In this arrange-

ment, transverse polarized light is incident on a

coupling prism at a specific angle extending an evanescent

Fig. 2 Optically resonant

photonic crystal biosensor from

Lee and Fauchet (2007).

Scanning electron microscopy

photograph of a typical device

used in these experiments and

schematic of the experimental

setup. A limit of detection

(LOD) on the order of 2.5 fg

(for bovine serum albumin) was

obtained using this device.

Copyright Optical Society of

America. Reproduced with

permission

38 Microfluid Nanofluid (2008) 4:33–52

123

(non-propagating mode) into the metal/dielectric interface.

Since the evanescent mode is non-propagating, under most

conditions the incident light is reflected from the surface

and very little loss in optical power is observed at the

detector. At a particular angle of incidence, hsp, a

momentum matching condition exists and a certain amount

of the incident energy is coupled into a surface plasmon

mode resulting in a reduction in the reflected power. The

incident angle for this plasmon excitation is described by

Eq. (1) below,

sin hsp ¼1

np

ffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi

emðkÞed

emðkÞ þ ed

s

ð1Þ

where np is the refractive index of the prism, em(k) is the

dielectric constant of the metal film which varies as a

function of the excitation wavelength, k, and ed which

represents the dielectric constants of the dielectric layer

(Homola et al. 1999a, b). Since the plasmon wave extends

only a few hundred nanometers into the dielectric layer

only the near field region is probed optically by the tech-

nique. At fixed wavelength em is constant and thus from

Eq. (1) it is apparent that hsp is only a function of the

dielectric constant in the region very near the interface.

For biosensing applications, this metal/dielectric inter-

face consists of a thin metal film deposited either directly

on the prism surface or an index matched slide and a

solution phase ‘‘sensing’’ (dielectric) layer. When the metal

layer (typically gold is used as it is inert to the atmosphere,

but can be modified by appropriate surface chemistry to

provide active sites for bonding of organic molecules) is

functionalized with a series of biorecognition agents, the

dielectric constant of the sensing layer is modified and a

shift in hsp is observed. As solution phase targets specific to

the immobilized biorecognition agents are introduced the

refractive index in the sensing layer increases further and

hsp again is affected. By monitoring the rate of and total

change in hsp, quantitative information regarding the

presence and concentration of the solution phase targets,

and bioaffinity constants for the target-probe binding

reaction can be obtained. A subtle variation on the above

geometry uses a grating coupler to excite a plasmon wave

at the interface (Unfricht et al. 2005).

3.3.2 Spectral SPR biosensing

A variation on the above approach, termed spectral SPR

(Dostalek et al. 2005; Otsuki et al. 2005; Yuk et al. 2005),

exploits the wavelength dependence of em to interrogate the

surface plasmon coupling conditions. The apparatus is

similar to the angular approach above, the difference being

that the experiment is conducted at fixed incident angle and

a tunable excitation source is used to sweep over the

wavelengths of interest as illustrated in Fig. 3b (alterna-

tively a broadband excitation source could be used and the

spectrum of the reflected light is recorded). It is not diffi-

cult to extract that since em is a function of k so long as the

wavelength scanning range is large enough that the right

hand side of Eq. (1) can be made to equal the left had side

at a particular resonant wavelength. Analogous to the

above, by monitoring the change in the resonant wave-

length with time information regarding the concentration

and bioaffinity properties of the target can be obtained.

PMT

Goniometer

I

Plasmon resonance

Prior to binding RXN

After binding RXN

Metal/Dielectric

Interface

Monochromaticlight source

PMT

I

Prior to binding RXN

After binding RXN

Tunablelightsource

Fixed position

Nano-particles

PMT

Tunablelightsource

λ

I

Prior to binding RXN

After binding RXN

θsp

(c)(b)(a)

θ λ

Fig. 3 Variations on surface

plasmon resonance (SPR)

detection. a Angular SPR,

b spectral SPR, and cnanoparticle or Local SPR.

Details on each of the above

techniques are provided in the

text

Microfluid Nanofluid (2008) 4:33–52 39

123

The spectral technique has several advantages over the

traditional angular approach which are usually stated as a

simplification of the experimental apparatus and increased

inherent sensitivity [since the wavelength dependence of em

is relatively weak (Homola et al. 1999a, b)]. More

importantly, however, the spectral approach lends itself to

imaging based data collection techniques (Otsuki et al.

2005) which are more amenable to multiplexing for high

throughput screening (as will be described below). Despite

these advantages the angular SPR approach is more

broadly exploited than the spectral approach due to added

expense of the tunable illumination source or the added

complication of extracting the reflected spectrum.

3.3.3 Local SPR biosensing

As an alternative to the traditional SPR biosensing

approaches outlined above, nanoparticle based SPR (or

local SPR, LSPR) has recently been developed for surface

phase geometries (Kelly et al. 2003). This approach relies

on coupling into a plasmon mode on the surface of a sub-

wavelength scale, surface immobilized, metallic nanopar-

ticle or nanostructure (Haes and Van Duyne 2002; Nath

and Chilkoti 2002; Yonzon et al. 2004) or nanohole array

(Tetz et al. 2006). Similar to the spectral approach above,

this results in a decrease in the transmitted power (or

increase in scattered power) at a specific resonant wave-

length which is strongly dependent on the environmental

dielectric conditions. The most significant advantage of this

approach is the simplicity of the arrangement (as shown in

Fig. 3c) and the facile extension to a highly multiplexed

architecture (details of how this is done will be expanded

on Sect. 3.3.5).

The most extensive characterizations and implementa-

tions of this architecture have been carried out by the Van

Duyne group (see a recent review by Zhao et al. 2006). The

well cited study by Haes and Van Duyne (2002) provides

details of the experimental apparatus and reports refractive

sensitivities on the order of Dn = 0.01. Implementation in

the context of the detection of biomarkers associated with

Alzheimer’s disease (Haes et al. 2005) using sera samples

has also been demonstrated. Readers are also referred to

another series of papers from this group (Haes et al. 2004;

Yonzon et al. 2004) which provides a comparison between

localized and propagating SPR biosensing. Interesting

recent examples from other groups include the work by

Kim et al. (2007) who developed an interesting LSPR

architecture based on what they referred to as a ‘‘gold

capped oxide nanostructure’’ comprising of oxide posts

with gold caps. Because the spacing between the structures

was relatively small the reflected light exhibited interfer-

ometric behavior that was highly sensitive to changes in the

thickness of the biomolecular layer. Detection limits in the

pM range (for oligonucleotides) were reported. The LOD

reported in some of the earlier papers using this approach

were much lower than those for traditional SPR. however.

recently Dahlin et al. (2006) described a technique for

improving the instrumentation and data analysis techniques

for this architecture in the end reporting sensitivities on the

order of 100 pg/cm2. In a different approach Kim et al.

(2006) compared the LSPR response of gold nanodot and

nanoring arrays (fabricated using imprint lithography),

examining the effect that nanostructure geometry has on

the internal sensitivity (e.g. decreasing ring width was

found to increase sensitivity).

One of the challenges with multiplexing such systems is

in the development of ordered and uniform metal nano-

structured arrays which are sufficiently well spaced to

avoid plasmon coupling effects (Haynes et al. 2003) but

sufficiently dense so that the aggregate signal can be

detected. Nanofabrication techniques such as focused ion

beam and electron beam lithography make precision fab-

rication of such arrays relatively easy but the serial nature

and general expense involved, strongly conflict with the

development of mass producible devices. As a result a

number of alternative parallel techniques have been

developed. One of the more commonly exploited tech-

niques is nanosphere lithography (Hulteen and Vanduyne

1995; Haynes and Van Duyne 2001) whereby a monolayer

of self-assembled spheres is assembled on a surface and a

metallic layer deposited over the array. After chemical

removal of the spheres, metal deposited through the voids

in the monolayer comprise the metallic nanoparticle array.

While the local quality of the array can be excellent, a

challenge remains in being able to self-assemble a defect

free sphere layer over a wide enough spatial distance to

extend this local order to larger area arrays. An alternative

technique based on nanoimprint lithography was demon-

strated by Kim et al. (2006).

3.3.4 Sensitivity and LOD of propagating

and localized SPR

To date the most sensitive propagating SPR measurements

are those reported by Tao et al. (1999) who demonstrated

the potential for detection resolution on the order of Dn =

10�8. In general, however, most systems operate with

resolutions on the order of Dn = 10�5. This results in

LODs which are typically on the order of 100–1,000 pg/

cm2 (Jung et al. 1998; Su et al. 2005). As with some of the

other devices above, the extreme sensitivity exhibited by

propagating SPR at least partially relates to the relatively

large interaction lengths of the plasmon field with the

surface bound analyte. As such a large number of binding

40 Microfluid Nanofluid (2008) 4:33–52

123

events are required in order to induce a relatively small

aggregate change in refractive index and rare target iden-

tification becomes more difficult. The ultimate limitation of

such systems is that the achievable spatial resolution is

limited by the decay length of the excited plasmons (Berger

et al. 1994). For most setups this is on the order of 20 lm

and we can use this to estimate the potentially achievable

LOD as being on the order of 0.4 and 4 fg (i.e. 0.4 · 10�15

and 4 · 10�15 g). It is important to note, however that, as

discussed by Huang et al. (2007), the propagation distance

of the plasmon can be substantially reduced by optimizing

the interrogation wavelength and thus there may be room to

push these values farther down.

As mentioned in the introductory paragraphs, the

strength of emerging nanotechnologies is not necessarily

that the overall internal sensitivity to bulk measures is

much greater but that the surface area that is probed is

much smaller. McFarland and Van Duyne (2003) have

reported zeptomole level sensitivity using single silver

nanoparticles sensors. A darkfield image of the silver

nanoparticles used in these experiments is shown in Fig. 4.

As expanded on in the McFarland and Van Duyne (2003)

work, a paper by Riboh et al. (2003) demonstrated wave-

lengths shifts on the order of 38 nm for adsorption of

*100 antibiotin molecules. Extrapolating then it can be

argued that assuming a molecular mass of 150 kDa for

antibiotin the total mass per particle is on the order of

25 ag (25 · 10�18 g). Assuming a 1 nm wavelength reso-

lution the LOD should be on the order of 0.6 ag.

3.3.5 Multiplexing SPR biosensing techniques

Surface plasmon resonance imaging (Rothenhausler and

Knoll 1988) is a relatively simple technique by which the

above measurements can be multiplexed to monitor 2D

arrays (Homola et al. 2005). In essence the technique

involves exciting a relatively broad area of the sample

surface which has been pre-arrayed with a series biorec-

ognition sites. For the propagating SPR case the reflected

light is then imaged on, typically, a CCD camera (in place

of the PMT shown in Fig. 3). The use of 2D imaging

makes it difficult to extract the optical adsorption as a

function of incident angle (Fig. 3a). As such the mea-

surement is typically made at a fixed wavelength and by

observing the changes in spatial pattern of the reflected

intensity as the peak adsorption angle drifts closer to or

farther away from the observation angle (Jordan et al.

1997; Nelson et al. 1999, 2001). Numerous examples of the

application of this technique exist in the literature as

applied to nucleic acid hybridization (Nelson et al. 2001)

and detection of low molecular weight protein biomarkers

(Lee et al. 2006).

While this method is relatively easy to setup and make

kinetic measurements with the relatively broad adsorption

peaks obtained from SPR measurements can reduce the

overall sensitivity (i.e. the percent change in reflectivity in

the case of a positive binding event can be quite small). To

improve the overall sensitivity of the technique Fang et al.

(2006) demonstrated a nanoparticle enhanced technique for

low level and multiplexed detection of microRNA. A

similar amplification method was used by Li et al. (2006)

for single nucleotide polymorphism genotyping. Shu-

maker-Parry and Campbell (2004) introduced an apparatus

that allowed the extraction of angular data in an imaging

format. In that work 1.2 ng/cm2 sensitivity was reported

and which can be equated to a 500 fg LOD for a 200 lm

spot size. As an alternative Otsuki et al. (2005) demon-

strated the use of spectral SPR imaging whereby images

are captured as a function of wavelength rather than

angular position. The advantage of this approach is that

eliminates the need for a rotating experimental setup, but

increases the complexity of the apparatus as either a

tunable laser source or monochromator is required.

To improve the LOD, imaging techniques based on

localized SPR are also under development. As an example

Endo et al. (2006) demonstrated an interesting core-shell

nanoparticle imaging technique applied to the monitoring

of antibody–antigen reactions. Using this technique as

many as 300 separate reaction sites were monitored. The

advantage of such techniques is the relative simplicity of

the optical setup (a simple co-linear transmission setup can

be used, rather than requiring a precise angular alignment)

and the potential to achieve the single nanostructure LOD

Fig. 4 Dark-field optical image of Ag nanoparticles used as single

particle nanosensors in McFarland and Van Duyne (2003). The field

of view in this image is approximately 130 lm · 170 lm. Copyright

Optical Society of America. Reproduced with permission

Microfluid Nanofluid (2008) 4:33–52 41

123

as described in Sect. 3.3.4. One of the major remaining

challenges is in the development of techniques for inde-

pendently functionalizing each nanostructure.

On the commercial side SPR biosensor developers have

introduced multiplexed versions of their existing platforms,

namely the Spreeta from Texas Instruments (Spangler et al.

2001; Chinowsky et al. 2003) and FlexChip from BIA-

CORE (Wassaf et al. 2006). Of those that which contains

the highest degree of parallelity is the FlexCHIP platform.

An example implementation of which for antibody arrays

is shown in Fig. 5, from Usui-Aoki et al. (2005). This chip

uses a grating coupling mechanisms and a mechanical

scanning technique to image as many as 400 reactions over

the course of 3 h (which is still orders of magnitude lower

than microarray based assays). In these multiplexed sys-

tems, however, the sensitivity is reduced to that available

from a standard microarray (see Usui-Aoki et al. 2005). As

such there remains a significant need for the development

of platforms which can maintain the extreme sensitivity

available from traditional SPR while pushing the limits of

parallelity into the 1,000s or 10s of 1,000s of binding

reactions.

4 Nanoscale electrical biosensors

In this section, we focus on the use of single dimensional

electrical nanostructures for biosensing. In the first section,

we provide a general overview followed by detailed dis-

cussion of the advantages of semiconductor nanowire and

carbon nanotube approaches. The second section addresses

issues relating to device assembly and multiplexing

architectures.

4.1 Single dimension electrical nanostructures

One-dimensional nanostructures (Xia et al. 2003; Lieber

and Wang 2007), such as carbon nanotubes and semicon-

ductor or polymer nanowires, represent an interesting and

relatively new paradigm for biochemical sensing (Cui et al.

2001). Though a number of different implementation

architectures exist, those which have been the most suc-

cessful mimic the basic field effect transistor, FET

(Cunningham 1998). In this arrangement, a single nanowire

(or nanotube) is placed between two lithographically pat-

terned microscale contact pads which are referred to as the

source and drain respectively. It is possible to functionalize

the nanowire with the appropriate capture probes prior to

assembly into the FET structure, however, due to the

complexity involved in the assembly process, it can be

easier to first assemble the wires into the device structure

and then functionalize them using traditional (e.g. spotting)

techniques. During sensing, the conductivity of the wire is

measured by monitoring the current flow between the

electrically excited contacts. When binding of a charged

species occurs, the local charge field is modified leading to

an accumulation or depletion of carriers in the nanowire,

reflected by a change in its conductivity (analogous to

applying a gate voltage). An example of this working

principal is shown in Fig. 6 for the nanowire virus

detection device presented by Patolsky et al. (2004). The

Fig. 5 Surface plasmon resonance antibody array chip from Usui-

Aoki et al. (2005). Antibody microarray layout. SPR signals from this

array were measured by a FLEXCHIPTM Kinetic Analysis System. aLayout of the array containing 400 reaction sites which could be

monitored in parallel. Array details are available in the

aforementioned reference. b Overview of the affinity chip where all

reaction sites are located in a 1 cm · 1 cm. A flow cell with a volume

of 47 lL is used to transport the sample over the array. cVisualization of immobilized antibodies. Copyright Wiley-VCH

Verlag GmbH & Co KGaA. Reproduced with permission

42 Microfluid Nanofluid (2008) 4:33–52

123

advantage of incorporating a nanowire into the FET is that

its cross-sectional area is of the same spatial scale as the

charge field surrounding the bound molecule and thus a

small number of close proximity targets can have signifi-

cant impact on the measured conductivity. The result is an

extremely sensitive device, capable of obtaining both static

and dynamic information through a very simple electrical

measurement.

4.1.1 Semiconductor nanowire based devices

Semiconductor nanowires have two main advantages over

carbon nanotubes when used in biosensing devices. First,

the material properties can be more precisely controlled by

manipulating the conditions during synthesis and using

well-developed semiconductor doping techniques. Second,

the native oxide layer that forms on the outside of silicon

nanowires allows the use of a broad class of already well-

developed functionalization and blocking chemistries.

Generally speaking, nanowire devices represent a suitable

technology for liquid phase measurements, so long as the

ionic strength of the sample buffer is relatively low. The

reason for this is that the spatial extent to which the charge

field of a bound target can expect to interact with the

nanowire is limited by the Debye screening length (Cheng

et al. 2006). This length is strongly dependent on the ionic

strength of the transport buffer with lower ionic strength

buffers yielding longer screening lengths (Hunter 1981).

Physiological systems such as human sera have relatively

high ionic strengths leading to screening lengths on the

order of 1 nm. As such, most devices require some dilution

of the physiological system in order to increase the Debye

length and detect binding events which may occur a fair

distance from the wire itself. Given this limitation, archi-

tectures have been used for a number of label free sensing

applications (see Li et al. 2004; Wang et al. 2005)

including the detection of single viruses (Patolsky et al.

2004). Of particular note is the work of Hahm and Lieber

(2004) who demonstrated that a time dependent conduc-

tance increase enabled the identification of fully

complementary versus mismatched DNA samples. For

more detailed information, biosensing applications are

covered in a number of broader reviews of semiconductor

nanowire devices (Bauer et al. 2004; Patolsky et al. 2006;

Wanekaya et al. 2006; Wang et al. 2007).

4.1.2 Carbon nanotube based devices

The competitive technology to the above is carbon nano-

tube field effect transistors (NT-FETs). Comprehensive

reviews are available for the synthesis and sensing of

carbon nanotube architectures (see Wang 2005; Sinha et al.

2006; Mahar et al. 2007). The two main advantages single

walled carbon nanotubes provide over silicon nanowires

are higher electron mobilities (McEuen et al. 2002) and

diameters in the sub nanometer range (Kong et al. 1998).

This makes it possible to, in principle, detect lower charge

densities. In addition to the limitations discussed above for

nanowires, an additional limitation of existing nanotube

devices is that the synthesis cannot be as precisely con-

trolled as with nanowire devices. Recent advancements,

however, enable oriented long single walled structures with

nearly homogeneous mechanical and electrical properties

(Huang et al. 2004) which may serve to solve this problem.

One of the first effective sensing architectures com-

prised of NT-FETs was presented by Kong et al. (2000),

where they were able to detect toxic gaseous molecules in

NH3 and NO2 down to concentrations as low as 1 ppm

with low response times (around 5 min) at room temper-

ature, which constituted a huge leap over the current

technology (Shimizu and Egashira 1999). It was also found

that an electrolyte gate is roughly an order of magnitude

Fig. 6 Technique for semiconductor nanowire based detection of

single viruses from Patolsky et al. (2004). In this figure, two nanowire

devices (labeled 1 and 2) are functionalized with different antibodies.

The antibodies on the second wire are specific to the target virus.

When introduced, the virus binds to the second nanowire and a

change in the conductance is observed and taken to be indicative of

the presence of the virus in solution. When the virus unbinds from the

nanowire, the conductance returns to its original value. Figure is

Copyright 2004 National Academy of Sciences, USA

Microfluid Nanofluid (2008) 4:33–52 43

123

more effective in modulating the conductance of a nano-

tube when compared to a standard silicon back gate

referenced to vacuum, making NT-FETs ideal for biolog-

ical applications (Rosenblatt et al. 2002). Likely the first

NT-FET displaying real time biological and pH sensing

capabilities was established by (Besteman et al. 2003) in

which they were able to measure glucose oxidase activity

by measuring changes in nanotube conductance. Part of

this breakthrough came as a result of using an effective

linker molecule between nanotube and the analyte. Sub-

sequent research by Larrimore et al. (2006) displayed how

nanotubes can function as ultra sensitive reference elec-

trodes to redox reactions, effectively giving rise to

nanoscale electrochemical cells. In parallel, single stranded

DNA was immobilized on aligned carbon nanotubes by He

and Dai (2004), who opened the route for sequence specific

DNA diagnosis using NT-FETs. In that work cyclic vol-

tammetry (Bard 2001) was used to analyze the binding

activity of two sets of complementary strands. NT-FET

sensors also present novel avenues for proteomic research.

Boussaad et al. (2003) were able to detect as few as 20

molecules of cythochrome c (*0.5 ag) adsorbed onto

individual nanotubes. In addition, Chen et al. (2004) were

able to accurately characterize protein binding and inhibi-

tion down to 100 nM solution phase concentration range.

4.2 Assembly and multiplexing techniques

Although nanowire fabrication and synthesis techniques

are very well developed (Wang 2003), one of the more

challenging aspect of incorporating them into a multi-

plexed sensing architecture has classically been the lack of

a simple technique for assembling the wires into a device

structure with the desired electrical connectivity. Typically

‘‘hybrid’’ assembly approaches are used which involve

lithographically patterning the electrical infrastructure and

then using an active assembly procedure to attract and

position the 1D nanostructure. Assembly techniques based

on a number of different mechanisms have been developed

include magnetic (Hangarter and Myung 2005), dielec-

trophoretic (Smith et al. 2000; Evoy et al. 2004) and fluidic

(Huang et al. 2001) positioning. The advantage of an active

approach (the former two) is that it enables one to, in

principle, capture different types (e.g. p-doped vs. n-doped

for nanowires, metallic vs. dielectric for nanotubes) or even

differently functionalized nanowires at specific locations,

thereby increasing the flexibility of the device. The chal-

lenge with this approach is the same as with most directed

assembly procedures in that although they may be auton-

omous, the amount of time required to complete a scaled

up assembly process (with say 10 s of thousands of reac-

tion sites) quickly becomes excessive. Passive assembly

approaches can be much quicker, but lack the positioning

specificity.

At present the most highly developed multiplexed

sensing architecture based on these ‘‘hybrid’’ assembly

procedures is that presented by the Lieber group (Zheng

et al. 2005) based on their pioneering semiconductor

nanowire studies (Cui and Lieber 2001; Cui et al. 2001;

Huang et al. 2001). The architecture as developed com-

prises a 1D array of approximately 200 electrically

independent nanowire devices and was used to perform

low-level detection of a series of serum-borne cancer

antigens. Very low solution level LODs were reported, on

the order of 0.9 pg/mL. Antibody attachment was carried

out using a well-developed functionalization chemistry

with traditional microarray spotting techniques used to

discriminate between different wires. Though only dem-

onstrated in 1D, an extension to 2D should be a relatively

simple matter simply involving slightly more involved

onboard electronics. To avoid the use of these relatively

complex assembly procedures, there is significant interest

in being able to develop nanowire sensing platforms

compatible with traditional CMOS fabrication. Such an

approach was recently demonstrated by Stern et al. (2007)

where nanowires were fabricated as part of a CMOS pro-

cess using ultra-thin SOI wafers and enabled through a

unique tetramethylammonium hydroxide wet etching pro-

cedure (traditional RIE type etches were observed to

degrade sensing performance). A LOD on the order of

100 fM (solution phase antibody concentration) was

reported.

5 Nanoscale mechanical biosensors

The next broad class of nanoscopic biosensor we will cover

here are those which exploit mechanical effects. We sep-

arate the discussion into two sections, the first focusing on

cantilever based approaches and the second emphasizing

acoustic techniques.

5.1 Cantilever based devices

The high mechanical quality factors associated with reso-

nant micro- and nano-mechanical systems (Craighead

2000; Majumdar 2002; Ekinci et al. 2004; Ekinci and

Roukes 2005) have proven useful in a number of applica-

tion areas ranging from signal processing (Shim et al.

2007), to electrometry (Cleland and Roukes 1998), to, as is

most relevant to this review, mass sensing. Recent works

have demonstrated the ability of doubly clamped very

high-frequency nanomechanical cantilever systems to

accurately weigh as little as a handful of xenon atom (Yang

44 Microfluid Nanofluid (2008) 4:33–52

123

et al. 2006) with a noise floor of *20 zg (LOD) and res-

olution of 7 zg or 7 · 10�21 g. As label free biosensors

such systems have been applied to the detection of viruses

(Ilic et al. 2004a, b; Gupta et al. 2006), single cells (Ilic

et al. 2001) and individual strands of DNA (Ilic et al.

2005). In these ‘‘dynamic mode’’ systems the amount of

bound mass is typically determined by observing changes

in the resonant frequency of the oscillator. Cantilever

readout is often done optically (Ilic et al. 2000) in which

case spatial multiplexing is relatively easy. The limitation

being that the cantilevers must be sufficiently well spaced

that they can be independently functionalized. While the

LOD of these biosensor can be as low as a few attograms

(Ilic et al. 2004a, b), see Fig. 7 for SEM images of the

devices used in that work, such measurements must nor-

mally be made in reasonably high vacuum environments

(or at least in air) due to the effect of viscous damping on

the mechanical oscillations. As a result, they must usually

be removed from the aqueous or physiological sensing

environment (the exception being the detection of air-borne

pathogens) before detection can be performed.

To avoid this problem several groups have developed

‘‘static mode’’ deflection based systems (Wu et al. 2001;

Zhang et al. 2006a, b), whereby the bending of a

mechanical structure is measured in response to binding

events along one side of the cantilever. Detection of

solution phase concentrations are on the order of pM,

which is roughly equivalent to that achievable using flo-

rescent labeling, have been reported. The LOD of these

devices tends to be lower since binding occurs along an

entire surface of the cantilever (which is on the order of

500 lm long · 100 lm wide). An alternative solution to

this problem has been presented by Burg and Manalis

(2003) who demonstrated the integration of microfluidic

channels inside a vibrating cantilever beam. The advantage

of this is that the cantilever itself can vibrate in a vacuum

environment, while the binding sites simultaneously

remain in contact with the liquid. An integrated version of

the device was recently reported (Burg et al. 2006) to have

an LOD on the order of 10�9 g/cm2 with eight indepen-

dently addressable sensor sites. With their reported surface

area (53,000 lm2) this yields an absolute LOD on the order

of 0.5 pg. Of additional note is a recent work by Braun

et al. (2005) which proposed a direct solution to the

vibrating cantilever in a physiological environment prob-

lem, analytically demonstrating than and LOD of 25 fg

should be possible. To our knowledge, however, this has

yet to be demonstrated experimentally.

5.2 Acoustic biosensors

Acoustic devices form a second class of mechanical bio-

sensor, the most common implementations of which

exploit bulk acoustic waves, BAW, such as the quartz

crystal microbalance (Marx 2003). Analogous to the above,

these devices measure changes in the resonant frequency of

a piezoelectric crystal in response to changes in surface

adsorbed mass. Limits of detection of such devices are on

the order of ng/cm2 and recent lithographically fabricated

devices (Rabe et al. 2003; Zhang and Kim 2005) have

demonstrated on-chip multiplexing with balance sizes on

the order of a few 10 s of micrometers in radius. A variant

on these BAW techniques are Surface Acoustic Wave,

SAW, devices which exploit shear surface waves (Josse

et al. 2001; Lange et al. 2006). Sensitivities are typically of

the same order (Kalantar-Zadeh et al. 2003) as QCM type

devices. Although the ultimate size of the interrogation

area is limited by the required spatial distance between

excitation electrodes, some recent works have discussed

techniques for exploit nanostructures to increase the overall

sensitivity (Rao and Zhang 2006). For more information

readers are referred to a recent review article by Lucklum

and Hauptmann (2006).

6 Homogeneous (solution) phase biosensors

Unlike the devices described above, mobile or solution

phase sensors typically incorporate functionalized nano-

particles, which have been suspended in the sensing

volume to act as both the binding and detection platform.

Fig. 7 Nanomechanical resonators for ultralow LOD mass sensors

from Ilic et al. (2004a, b). Oblique-angle SEM micrographs of a–dcantilevers (scale bar = 5 lm) and e–h bridge oscillators (scale

bar = 2 lm). The diameters of the Au pads were 50, 100, 200, and

400 nm, from left to right. Reused with permission from (Ilic et al.

2004a, b). Copyright 2004, American Institute of Physics

Microfluid Nanofluid (2008) 4:33–52 45

123

The advantage of such techniques is the ability to con-

centrate sensing units within a volume region as opposed to

a surface, which allows for faster and more efficient

binding of targets (as the diffusive transport length scale is

reduced). Such systems are also very amenable to in vivo

sensing applications and have the potential for very simple

feedback mechanisms. As will be discussed below, the

current challenges in this area revolve around the synthesis

of bio-compatible nanoparticles for these in vivo applica-

tions and the differentiation of different nanoparticles

within a single solution for multiplexed analysis.

6.1 Solution phase SPR

Similar LSPR techniques to those described in Sect. 3.3.3

are also popular in solution phase systems (Storhoff et al.

1998) since they have the ability to provide a very simple

colorimetric feedback (i.e. a change in the local dielectric

conditions surrounding a solution phase nanoparticle

results in a change in the peak scattered light which is

interpreted as a change in solution color by the observer).

Examples of implementation include detection of glucose

(Aslan et al. 2004), protein conformation changes (Chah

et al. 2005), specific sequences of genomic DNA (Li and

Rothberg 2004) and cholera toxin (Schofield et al. 2007).

As alluded to above, however, the relatively broad

absorption spectrum associated with plasmon coupling and

the inability to spatially localize the probes complicates

multiplexed detection. Yu and Irudayaraj (2007) reported

the use of functionalized gold nanorods for the detection of

shifts in plasmon waves due to molecular binding of pro-

teins on the surface of the nanorods. Use of controlled

aggregation of the gold nanorods allows for magnified

shifts in the plasmon peak. The plasmon bands show sig-

nificant sensitivity to the aspect ratio of the nanorods,

allowing for differentiation and increases the potential for

multiplexed analysis. In addition, spectra of nanorods with

multiple detection sites were shown to have differentiation

for single, double, and triple binding of targets. For more

information and a more in depth comparison with surface

phase techniques readers are referred to Haes et al. (2004).

6.2 Encoded quantum dots

Quantum dots (QDs) are very small (2–10 nm, not inclu-

sive of solubility ligands and functionalization chemistries)

semiconductor nanostructures which confine electrons to

discrete energy levels. The spacing between these energy

levels is strictly governed by the quantum confinement

conditions enforced by the dot size. As such the wave-

length at which a photon is emitted as the electron falls

from a higher to a lower energy state can be precisely

controlled through careful QD synthesis. Advantages of

QDs over most florescent dyes include: a wide absorption

spectrum coupled with a very narrow emission spectrum, a

high quantum yield, and a resistance to photobleaching and

optical or chemical degradation (Caruso 2004).

One of the most interesting approaches to multiplexing

solution phase systems uses a QD encoding scheme. This

approach involves the use of micro- or nanoparticles that

contain within themselves a series of embedded QDs of

different sizes. As such the emission spectrum of such a

particle comprises the summation of the emissions of the

QD particles embedded within. A particle then is designed

to have a specific colorimetric code which differs from

other particles made with a different ratio of colored QDs.

Quantum dots are particularly useful for this application as

they can all be excited with the same light source. Han

et al. (2001) used CdSe/ZnS QDs to demonstrate the

encoding of polymeric microbeads using a three-color

scheme. Cao et al. (2006) was able to encapsulate QDs

within a silica shell nanoparticle for improved biocom-

patibility. In both studies, the nanoparticles were

functionalized for the detection of fluorescently tagged

DNA. One of the advantages to this is the ability to

simultaneously detect both the coding and target spectrum.

Although a separate peak can be detected for a target sig-

nal, there is some overlap between the target FITC signal

and blue QD emissions. A review article by Medintz et al.

(2005) discusses many available functionalization methods

for labeling and sensing applications.

An example of research in multiplexed assays using

encoded QDs exploits flow cytometry for a fast, efficient,

and low-cost spectral analysis (Xu et al. 2003). Gao and

Nie (2004) demonstrated a system using flow cytometry

capable of reading 1,000 encoded particles per second.

Ideally, a sample would be mixed with a suspension of QD

embedded nanoparticles to bind targets to receptors. The

mixture would then be fed to a flow cytometry system

where the coding and target signals could be recorded for

later analysis.

6.3 Alternative solution phase approaches

Before closing this section we briefly mention a couple of

alternative approaches to multiplexed solution phase

sensing which do not fit into the previous two categories

but are important for comprehensiveness. Wang and Tan

(2006) recently developed a coding scheme based upon

fluorescence resonance energy transfer (FRET) using

varying ratios of inorganic and organic dyes. This partic-

ular method is valuable to biological in vivo applications

where toxicity concerns may prohibit the use of QDs.

46 Microfluid Nanofluid (2008) 4:33–52

123

Pregibon et al. (2007) demonstrated the use of microflui-

dics combined with photolithography to create graphically

encoded polymer nanoparticles that contain regions for a

grid identification system and analyte binding. The unique

identification scheme presented does not rely upon wave-

length or intensity based differentiation. Furthermore, these

particles due to the photolithographic synthesis method

have a potentially more stable and efficient yield.

7 Summary and conclusions

The previous four sections have provided a cross-sectional

overview of the state of the art in nanobiosensors, with

particular attention paid to how well each technology meets

the broad application requirements as outlined in Sect. 2.

Before closing this review we will attempt to condense as

much of that information as possible down into a few short

summary statements.

As was originally mentioned in section two and referred

to throughout this review, the primary advantage of

nanoscopic biosensors in not necessarily that they exhibit

higher internal sensitivity to bulk measures, such as

refractive index changes, but rather that the surface area or

volume which is probed is much smaller. Given that most

of the applications described above are concerned with

pushing the limits of how few molecules can be detected in

a given volume of solution, nanobiosensors are inherently

useful so long as the entire platform (inclusive of the

biosensor and associated micro-/nanofluidics) are designed

in such a way that the entire sample volume can be inter-

rogated by the sensor. With this in mind, to our knowledge

mechanically resonant devices have demonstrated the

lowest limits of detection, so long as the measurements can

be made in an environment where viscous damping is

minimized (usually vacuum but also potentially in air). In

liquid environments, where viscous damping is high, the

1D nanostructure electrical detection technologies have, to

date, best demonstrated the ability to push the limits of

sensitivity. Since most biomolecular detection strategies

involve the ‘‘chaining’’ together of a series of molecules

(namely a linker, probe and finally target) to perform the

measurement, the spatial limitations on the charge field

disturbance in moderate ionic strength solutions can in

some cases prove to be a significant limitation. As men-

tioned above, reducing the run buffer ionic strength can

help to increase the double layer thickness, but can also

serve to impede binding due to enhanced electrostatic

repulsion. Optical devices have a similar spatial probing

limitation, governed by the thickness of the evanescent

field. This tends to be somewhat less restrictive since the

evanescent field can be as thick as a few hundred nanom-

eters (Saleh and Teich 1991). The tradeoff is that the

thicker the evanescent field, the more the exposed optical

energy is diluted and the greater the probed volume. As

such a greater amount of bound mass is required to produce

a measurable change in the transduction signal. For sys-

tems like the zero-dimensional LSPR nanoparticles shown

in Fig. 4, the total surface area is so small that this is more

than compensated for.

Surface plasmon resonance imaging techniques as

described in Sect. 3.3.5 almost certainly represent the most

well developed and cheapest technique for performing

label free, highly multiplexed, biosensing. The simplest

implementations, however, lack the LOD strength of the

nanosensor technologies. As such platforms which incor-

porate the LOD advantages of the LSPR techniques while

maintaining this level of multiplexing are likely to be very

successful. The advanced optical devices such as the

photonic structures described in Sects. 3.1 and 3.2 have the

advantage of being able to operate in a planar format

(meaning the excitation source, device, fluidics, and

detector can all be in the same plane) make such devices

useful for more chip-based devices. To date, however, the

LOD and measurement parallelity of these devices has not

been as well developed as some other techniques. An

advantage of such devices is that the fabrication is not

particularly difficult, typically requiring only single layer

lithography. One-dimensional nanostructure arrays are

particularly promising given the demonstrated sensitivity

(subject to the limitations described above) and the inher-

ent ease of using an electrical measurement scheme. The

challenges in terms of assembly of nanowires into a device

structure are somewhat of a bottle neck at present, how-

ever, as the semiconductor industry continues to push the

limits of lithography, direct patterning will become more

and more feasible (beyond what has already been

demonstrated).

Another major challenge to the further development of

particle nanosensor technology is one that is not often

addressed in the literature, namely the ability to target and

confine chemical functionalization to the sensor site itself.

For example, sensors which probe only a small area but

require functionalization of a larger one (say due to the

use of traditional spotting techniques) do not really

improve the LOD as binding occurs everywhere and not

preferentially at the sensor site. In cases where the

background surface chemistry is much different than the

sensor itself (say a gold nanoparticle on a silica surface)

specific functionalization can be relatively easy. It is more

difficult for devices where the surface of the sensor itself

is not much different than the background surface. An

example of such a case would be silicon on insulator

optical devices (SOI) where it is difficult to find chem-

istries which will work on silicon but not the underside

oxide layer.

Microfluid Nanofluid (2008) 4:33–52 47

123

Acknowledgments The authors of this article are thankful for the

financial support of the National Science Foundation under grant

number CBET-0529045, the National Institutes of Health—National

Institute of Biomedical Imaging and Bioengineering under grant

number R21EB007031, the Biomolecular Devices and Analysis

program of the Nanobiotechnology Center at Cornell University, and

the Defense Advanced Research Projects Agency Young Investigator

Program.

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